The field of the invention is systems and methods for magnetic resonance imaging (“MRI”). More particularly, the invention relates to systems and methods for chemical exchange saturation transfer (“CEST”) MRI.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the nuclei in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) that is in the x-y plane and that is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped,” into the x-y plane to produce a net transverse magnetic moment Mxy. A signal is emitted by the excited nuclei or “spins,” after the excitation signal B1 is terminated, and this signal may be received and processed to form an image.
When utilizing these “MR” signals to produce images, magnetic field gradients (Gx, Gy, and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received MR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
The measurement cycle used to acquire each MR signal is performed under the direction of a pulse sequence produced by a pulse sequencer. Clinically available MRI systems store a library of such pulse sequences that can be prescribed to meet the needs of many different clinical applications. Research MRI systems include a library of clinically-proven pulse sequences and they also enable the development of new pulse sequences.
The MR signals acquired with an MRI system are signal samples of the subject of the examination in Fourier space, or what is often referred to in the art as “k-space.” Each MR measurement cycle, or pulse sequence, typically samples a portion of k-space along a sampling trajectory characteristic of that pulse sequence. Most pulse sequences sample k-space in a raster scan-like pattern sometimes referred to as a “spin-warp,” a “Fourier,” a “rectilinear,” or a “Cartesian” scan. The spin-warp scan technique employs a variable amplitude phase encoding magnetic field gradient pulse prior to the acquisition of MR spin-echo signals to phase encode spatial information in the direction of this gradient. In a two-dimensional implementation (“2DFT”), for example, spatial information is encoded in one direction by applying a phase encoding gradient, Gy, along that direction, and then a spin-echo signal is acquired in the presence of a readout magnetic field gradient, Gx, in a direction orthogonal to the phase encoding direction. The readout gradient present during the spin-echo acquisition encodes spatial information in the orthogonal direction. In a typical 2DFT pulse sequence, the magnitude of the phase encoding gradient pulse, Gy, is incremented, ΔGy, in the sequence of measurement cycles, or “views” that are acquired during the scan to produce a set of k-space MR data from which an entire image can be reconstructed.
There are many other k-space sampling patterns used by MRI systems. These include “radial,” or “projection reconstruction,” scans in which k-space is sampled as a set of radial sampling trajectories extending from the center of k-space. The pulse sequences for a radial scan are characterized by the lack of a phase encoding gradient and the presence of a readout gradient that changes direction from one pulse sequence view to the next. There are also many k-space sampling methods that are closely related to the radial scan and that sample along a curved k-space sampling trajectory rather than the straight line radial trajectory.
An image is reconstructed from the acquired k-space data by transforming the k-space data set to an image space data set. There are many different methods for performing this task and the method used is often determined by the technique used to acquire the k-space data. With a Cartesian grid of k-space data that results from a 2D or 3D spin-warp acquisition, for example, the most common reconstruction method used is an inverse Fourier transformation (“2DFT” or “3DFT”) along each of the 2 or 3 axes of the data set. With a radial k-space data set and its variations, the most common reconstruction method includes “regridding” the k-space samples to create a Cartesian grid of k-space samples and then performing a 2DFT or 3DFT on the regridded k-space data set. In the alternative, a radial k-space data set can also be transformed to Radon space by performing a 1DFT of each radial projection view and then transforming the Radon space data set to image space by performing a filtered backprojection.
Measuring the exchange of magnetization between molecules with nuclear magnetic resonance can provide unique information about the chemical and molecular environment of samples or tissues. One type of such exchange measurement is broadly referred to in the field as magnetization transfer. This technique is capable of measuring the exchange of magnetization from spin species that have very short transverse relaxation times (“T2”). Because many different molecules have short T2, this technique is not particularly specific.
A second type of magnetization exchange occurs between water protons and a molecule with long enough T2 that its difference in frequency from water can be observed. Saturation of the magnetization from this molecule will generally decrease the measurable signal from water. This effect is generally referred to in the field as chemical exchange saturation transfer (“CEST”). Two different types of molecules can generate CEST effects: endogenous, or naturally occurring, molecules and exogenous contrast agents. In either instance, the molecules whose chemical exchange with water produces the CEST effect are generally referred to as a so-called “exchangeable protons.”
The CEST imaging method offers three advantages over traditional molecular MRI techniques. First, in some cases the molecules of interest within the subject can be directly detected. This feature mitigates the need for administering contrast agents to the subject. Second, the image contrast mechanism can be controlled with the RF pulses produced by the MRI system and, as such, can be turned on and off when desired. This control allows the location of specific molecules of interest to be detected by comparing images having the desired contrast present to those where it has been turned off. Lastly, the CEST imaging method is far more sensitive than traditional molecular MRI techniques, making it able to detect substantially low concentrations of given molecules.
A number of different molecular groups have been suggested for CEST studies. One such group are the amide protons. Amide protons are present in large numbers on peptides and proteins; therefore, amide proton CEST should be reflective of protein concentration in cells. However, other exchangeable protons are also targeted with CEST imaging methods. Exemplary exchangeable protons include those protons contained in hydroxyl and glycogen, as well as paramagnetic molecules in general.
The size of the CEST effect is determined by how quickly the protons exchange their magnetization with water. This exchange rate is believed to be determined by pH, so the CEST effect can also potentially provide information indicative of altered pH levels. Conventional amide proton CEST imaging methods are described, for example, by J. Zhou, et al., in “Using the Amide Proton Signals of Intracellular Proteins and Peptides to Detect pH Effects in MRI,” Nature Medicine, 2003; 9:1085-1090, and in U.S. Pat. No. 6,943,033. Imaging using amide proton transfer contrast has a number of potential applications. For example, low pH is indicative of ischemia and could be used in imaging of stroke and other ischemic diseases.
For amide proton CEST, and many other endogenous CEST applications, a major difficulty arises when attempting to avoid other sources of signal change when saturation of the labile proton line is performed. For example, off-resonance saturation can cause direct saturation of the nearby water line, as well as magnetization transfer effects resulting in saturation of broad molecular lines that exchange magnetization with water.
CEST experiments rely on the difference in CEST effect with the frequency of application. One common imaging approach is to compare an image with saturation applied at the frequency of the molecule of interest with another where the saturation is applied on the opposite side of the water frequency, that is applied at the negative of the first frequency. This is successful only if the position of the water line is known exactly and if the magnetization transfer effect is symmetric around the water line. While these situations might be achievable in vitro, neither are present in vivo, thereby resulting in significant errors when utilizing the aforementioned approach. An alternative is to perform CEST studies at many frequencies and fit the signal as a function of frequency to some model. In addition the fact that the appropriate model is typically not known, the long acquisition time needed to acquire images over many frequencies is a disadvantage. Even with the acquisition of many data points, amide proton CEST in normal tissue is very difficult to measure. CEST, and specifically amide proton transfer CEST, is not currently used in diagnostic applications due to the foregoing difficulties with the imaging method.
It would therefore be desirable to provide a method for chemical exchange saturation transfer (“CEST”) imaging in which errors resultant from off-resonance and magnetization transfer effects are substantially mitigated, and in which clinically satisfactory imaging times are achieved.